A Label-Free Electrochemical Impedimetric Immunosensor with Biotinylated-Antibody for SARS-CoV-2 Nucleoprotein Detection in Saliva

The timely detecting of SARS-CoV-2 coronavirus antigens for infection validation is an urgent request for COVID-19 pandemic control. This study constructed label-free electrochemical impedance spectroscopy (EIS)-based immunosensors based on gold nanostructured screen-printed carbon electrodes (AuNS/SPCEs) to detect the SARS-CoV-2 nucleocapsid protein (N-protein) in saliva. Using short-chain 3-mercaptopropionic acid (MPA) as a linker to covalently bond streptavidin (SA) and bovine serum albumin (BSA) for controlling the oriented immobilization of the biotinylated anti-N-protein antibody (BioAb) can offer a greater sensitivity, a lower limit of detection (LOD), and better reproducibility of immunosensors (defined as BioAb/SA-BSA/MPA/AuNS/SPCEs) than the antibody randomly immobilized immunosensors and the long-chain 11-mercaptoundecanoic acid (MUA)-modified immunosensors (BioAb/SA-BSA/MUA/AuNS/SPCEs). The BioAb/SA-BSA/MPA/AuNS/SPCE-based immunosensors presented good linearity from 0.01 ng/mL to 100 ng/mL and a low LOD of 6 pg/mL in a phosphate buffer solution (PBS) and PBS-diluted saliva. Moreover, the immunosensor exhibited little cross-activity with other viral antigens such as MERS-CoV N-protein, influenza A N-protein, influenza B N-protein, and SARS-CoV-2 spike protein, indicating the high specificity of the immunosensors. The disposable label-free EIS-based immunosensors have promising potential in facilitating the rapid and sensitive tests of saliva-based COVID-19 diagnostics.


Introduction
Since the coronavirus disease 2019  was declared by the World Health Organization (WHO) in 2020, the cumulative death toll is over six million to date [1]. The COVID-19 pandemic, caused by the severe acute respiratory syndrome coronavirus 2 (SARS-CoV-2), has induced a catastrophic impact on the human healthcare system and economy [2]. The symptoms of COVID-19, from mild to severe, including fever, cough, headache, loss of taste or smell, sore throat, and pulmonary infiltrates, may appear 2-14 days after exposure to the virus [3]. However, some infected patients are asymptomatic. These asymptomatic patients still have a great potential to spread the COVID-19 disease to other people. Although vaccination can reduce the risk of infection and severe symptoms of COVID-19 against viral transmission, the breakthrough infection and the asymptomatic infection still occur by an emerging variant virus [4]. Therefore, developing a rapid, sensitive, and accurate detection platform to diagnose COVID-19 carriers is crucially essential for preventing the disease from spreading at the early stages.
COVID-19 diagnostic testing can be classified into two categories [5][6][7][8][9]: the viral tests, including antigen tests and nucleic acid amplification tests (NAATs), and the antibody bilization strategies were compared to optimize the sensing properties. Moreover, the detecting ability of immunosensors in the N-protein-spiked saliva samples was explored in detail.

Immunosensor Preparation
The AuNS electrodeposition procedure of SPCEs refers to our previous studies [40][41][42]. Initially, the SPCEs were cleaned by a cyclic potential from 0 to 1.3 V for 20 cycles with a 0.1 V/s scanning rate and then oxidized at 2.0 V for 30 s in 0.1 M PBS (pH 7.0) with a three-electrode system using a Pt plate as the counter electrode and a commercial Ag/AgCl electrode as the RE. These procedures can increase the hydrophilicity and electron transfer rate of SPCE surfaces. Subsequently, the oxidized SPCEs were dipped in the 100 mM KCl-containing HAuCl 4 solution (8 mM, pH 2.0) for the two-step AuNS deposition. The first step was AuNP nucleation on the oxidized SPCEs using a cyclic potential from 0.5 to −0.5 V with a scan rate of 50 mV/s for seven cycles. The second step was AuNS formation on the AuNPs with a step potential at 0.62 V for 10 min. The AuNS/SPCEs were rinsed with double-distilled water to remove the free ions from the electrode surface for subsequent surface modification.
A small amount (10 µL) of aliquot of 10 mM MPA or 10 mM MUA was dripped on the AuNS/SPCEs at 30 • C for 1 h in an incubator to form a self-assembled monolayer (SAM) as a linker for the immobilization of SA and BSA. The unbound MPA molecules were removed using double-distilled water. The carboxyl group of the MPA/AuNS/SPCEs was activated by the EDC (30 mM)/NHS (30 mM) mixture-containing 50 mM MES solution (pH 4.6) for 1 h. After rinsing the electrodes with PBS, 10 µL aliquot of SA (150 µg/mL)/BSA (150 µg/mL) mixture was placed on the activated MPA/AuNS/SPCEs or MUA/AuNS/SPCEs for 1 h. After rinsing the electrodes with PBS, the 10 µL BioAb (100 µg/mL) solution was placed on the SA-BSA-immobilized electrodes for 1 h, then dipped in the 0.0025% Tween 20-containing PBS for 10 min to remove the unbound BioAb. The biotinylation recipe of BioAb followed the Thermo company's suggestion. Some (125 µL) aliquots of 2 mg/mL anti-SARS-CoV-2 N-protein Ab was mixed with 0.8 µL of 50 mg/mL EZ-Link Sulfo-NHS-LC-Biotin for 30 min at 25 • C, and then, 17.05 µL aliquot of 2.5 mM glycine was added to block the reaction for 1 h at 25 • C. Before use, the BioAb solu-tion was stored at 4 • C. The immunosensors were incubated with the concentration-varied N-protein samples prepared in PBS or ten times-diluted salivary solutions for 40 min and then rinsed with PBS. EIS estimated the change in the electrochemical properties of the electrode/solution interface. The preparing procedures of the immunosensors are shown in Scheme 1. the 10 μL BioAb (100 μg/mL) solution was placed on the SA-BSA-immobilized electrode for 1 h, then dipped in the 0.0025% Tween 20-containing PBS for 10 min to remove th unbound BioAb. The biotinylation recipe of BioAb followed the Thermo company's sug gestion. Some (125 μL) aliquots of 2 mg/mL anti-SARS-CoV-2 N-protein Ab was mixe with 0.8 μL of 50 mg/mL EZ-Link Sulfo-NHS-LC-Biotin for 30 min at 25 °C, and then, 17.0 μL aliquot of 2.5 mM glycine was added to block the reaction for 1 h at 25 °C. Before use the BioAb solution was stored at 4 °C. The immunosensors were incubated with the con centration-varied N-protein samples prepared in PBS or ten times-diluted salivary solu tions for 40 min and then rinsed with PBS. EIS estimated the change in the electrochemica properties of the electrode/solution interface. The preparing procedures of the im munosensors are shown in Scheme 1.

Electrochemical Measurements
An equimolar Fe(CN)6 3−/4− mixture (2.5 mM) in 10 mM PBS (pH 7.4) was used as mediator to estimate the electrochemical properties of the electrodes in each modificatio step and quantify the immunoreaction by an EIS workstation (MultiPalmSens4, PalmSen Holten, The Netherlands). Some (100 μL) aliquots of the mediator-containing PBS wer placed on the immunosensor surface for the EIS measurements. The EIS parameters wer set in 1-100 kHz at +0.1 V versus the SPCE-based pseudo-RE, added by a 5 mV amplitud sine wave. The impedance spectra and the equivalent circuit simulation were measure using the MultiTrace-4.4 software package (PalmSens).

Antibody Immobilization Strategies
The paratope orientation of immobilized Ab affects the immunoreaction efficiency o immunosensors [43]. Figure 1 shows the Nyquist plots obtained in each step of im munosensor preparation with different immobilization strategies and their immunoreac tion with 50 μL aliquot of 1 ng/mL N-protein. Figure 1a−c show the random Ab immob lization on a short-chain MPA SAM, the oriented Ab immobilization on a MPA SAM, an a long-chain MUA SAM, respectively. The inset of Figure 1a shows the impedance spectr measured at the bare AuNS/SPCEs, which exhibited a dominant linear region, implyin an apparent diffusion-controlled behavior, and a small semicircle region of kinetic contro attributed to the fast electron transfer rate of the Fe(CN)6 3−/4− mediator on the high conduc tive surface of AuNS/SPCEs [40−42]. Following MPA modification, the Nyquist plo showed a small linear part and a large semicircle part, implying a decreasing electro

Electrochemical Measurements
An equimolar Fe(CN) 6 3−/4− mixture (2.5 mM) in 10 mM PBS (pH 7.4) was used as a mediator to estimate the electrochemical properties of the electrodes in each modification step and quantify the immunoreaction by an EIS workstation (MultiPalmSens4, PalmSens, Holten, The Netherlands). Some (100 µL) aliquots of the mediator-containing PBS were placed on the immunosensor surface for the EIS measurements. The EIS parameters were set in 1-100 kHz at +0.1 V versus the SPCE-based pseudo-RE, added by a 5 mV amplitude sine wave. The impedance spectra and the equivalent circuit simulation were measured using the MultiTrace-4.4 software package (PalmSens).

Antibody Immobilization Strategies
The paratope orientation of immobilized Ab affects the immunoreaction efficiency of immunosensors [43]. Figure 1 shows the Nyquist plots obtained in each step of immunosensor preparation with different immobilization strategies and their immunoreaction with 50 µL aliquot of 1 ng/mL N-protein. Figure 1a−c show the random Ab immobilization on a short-chain MPA SAM, the oriented Ab immobilization on a MPA SAM, and a long-chain MUA SAM, respectively. The inset of Figure 1a shows the impedance spectra measured at the bare AuNS/SPCEs, which exhibited a dominant linear region, implying an apparent diffusion-controlled behavior, and a small semicircle region of kinetic control attributed to the fast electron transfer rate of the Fe(CN) 6 3−/4− mediator on the high conductive surface of AuNS/SPCEs [40][41][42]. Following MPA modification, the Nyquist plot showed a small linear part and a large semicircle part, implying a decreasing electron transfer rate. Moreover, the semicircle radius increased with the Ab modification and the 40-min immunoreaction of the N-protein. The phenomenon indicates that the impedance of the solution/electrode interfaces increased with the modification and the immunoreaction. The corresponding electric elements fitted by the modified Randles equivalent circuit are listed in Table 1. The modified Randles equivalent circuit, consisting of the solution resistance (R s ), the Warburg impedance (Z w ), the constant phase element (CPE), and the electron transfer resistance (R et ), was mentioned in our previous articles to explain the diffusive and kinetic behavior of the solution/electrode interface [40][41][42]. The mean error of all the fitting data was less than 0.3%. transfer rate. Moreover, the semicircle radius increased with the Ab modification and the 40-min immunoreaction of the N-protein. The phenomenon indicates that the impedance of the solution/electrode interfaces increased with the modification and the immunoreaction. The corresponding electric elements fitted by the modified Randles equivalent circuit are listed in Table 1. The modified Randles equivalent circuit, consisting of the solution resistance (Rs), the Warburg impedance (Zw), the constant phase element (CPE), and the electron transfer resistance (Ret), was mentioned in our previous articles to explain the diffusive and kinetic behavior of the solution/electrode interface [40−42]. The mean error of all the fitting data was less than 0.3%.    Figure 1b shows the Nyquist plots of AuNS/SPCEs, followed by MPA modification, SA-BSA immobilization, BioAb affinity attachment, and 1 ng/mL N-protein immunoreaction. The semicircle radius of the EIS plots increased with the step-by-step modification and the immunoreaction, indicating the increasing R et . It is worth noting that the EIS plot only presented a semicircle after the N-protein immunoreaction, implying a slow electron transfer rate to dominate the Faradaic reaction, with little diffusion-controlled behavior. In our previous studies [40][41][42], the 1R//C equivalent circuit, consisting of R s in a series with one parallel circuit comprising a R et and a CPE, was used to analyze the EIS spectrum of only the semicircle part. The corresponding circuit element values are statistically calculated as the mean ± standard deviation from three individual immunosensors in Table 1 Notably, the ∆R et /R et0 was 2.9 times larger than that obtained at the Ab/MPA/SPCEs, elucidating that the oriented immobilization of BioAb on the SA-BSA-modified electrodes had a higher immunoreacting efficiency than the random Ab immobilization. Many groups have demonstrated oriented Ab immobilization by using biotin-streptavidin affinity, protein A or G for the adsorption of the Ab Fc portion, or recombinant peptide tags for metal coordination with the promoted sensitivity of the immunosensors [42][43][44]. Lin et al. found that the EIS-based immunosensors with the oriented Ab immobilization on the protein A (PA, 100 µg/mL)-modified MPA/AuNS/SPCEs had a sensitivity larger than those with the random Ab immobilization [42]. Moreover, the PA (100 µg/mL)-to-BSA (100 µg/mL) ratio-modified MPA/AuNS/SPCEs could adsorb more antibodies and had a lower limit of detection (LOD) than the PA (100 µg/mL)-modified MPA/AuNS/SPCEs. Therefore, the study adopted the same concentration ratio (1:1) of SA to BSA to modify the MPA/AuNS/SPCEs. Furthermore, the MPA SAM is thin enough to produce a large CPE. The CPE value decreases significantly with the SA-BSA immobilization, Ab adsorption, and N-protein immunoreaction, attributed to layer-by-layer stacking with an increasing thickness of the modifying layer. Figure 1c shows the EIS plots with only the semicircle part after the MUA modification. Moreover, the semicircle radius of the MUA/AuNS/SPCEs was more prominent than that of the SA-BSA-and BioAb-modified electrodes. After being analyzed by the 1R//C equivalent circuit, the R et of the MUA/AuNS/SPCEs was much larger than that of the MPA/AuNS/SPCEs (shown in Table 1), resulting from the dense and long MUA SAM [41]. The high-coverage MUA COO − groups presented strong electrostatic repulsion for the negatively charged Fe(CN) 6 3−/4− . After EDC/NHS activation and SA-BSA immobilization, the R et of the SA-BSA/AuNS/SPCEs was smaller than that of the MUA/AuNS/SPCEs, attributed to the drastic decrease of the MUA COO − groups. Although the ∆R et (23.3 kΩ) of the BioAb/SA-BSA/MUA/SPCEs was much larger than that (1.43 kΩ) of the BioAb/SA-BSA/MPA/SPCEs after the 1 ng/mL N-protein immunoreaction, the ∆R et /R et0 (19.3%) was smaller than that (77.7%) of the BioAb/SA-BSA/MPA/SPCEs. The results elucidated that the MPA linker and SA-BSA layer combination can produce a more sensitive immunosensor, and the ∆R et /R et0 is more significant than the ∆R et for EIS signal compared between fabrication-varied immunosensors. Figure 1d shows the time-dependent immunoreaction curve. The ∆R et values measured at 50 min reached a saturated reaction. We wanted to compromise the time and response magnitude of the immunoreaction to select a 40-min immunoreaction to investigate other sensing properties.  , higher than that of the MPA-based immunosensors. This phenomenon was attributed to the high insulation and thickness of the MUA SAM to reduce the electric response in the Ab-antigen interface [41]. Moreover, the RSD of the calibration curve ranged from 0.4% to 3.1%, implying a high reproducibility of the MUA-based immunosensors.

Other Sensing Properties
The BioAb/SA-BSA/MPA/SPCE-based immunosensors were tested against different viral antigens, including the MERS-CoV N-protein, influenza A virus N-protein, influenza B virus N-protein, and SARS-CoV-2 S-protein, to investigate the immunosensor specificity. Figure 3 shows the ∆R et /R et0 response of each analyte in low (1 ng/mL) and high (100 ng/mL) concentrations. Compared to the ∆R et /R et0 values of 1 ng/mL (77.7%) and 100 ng/mL (146.2%) SARS-CoV-2 N-protein, the ∆R et /R et0 values of the other analytes ranged from 0.51% to 1.21% and from 0.58% to 1.84%, respectively; those are much smaller than the ∆R et /R et0 of SARS-CoV-2 N-protein. The signal-to-cross reactivity ratio of 1 ng/mL and 100 ng/mL SARS-CoV-2 N protein was larger than 64.2 and 79.5, respectively. The results indicated that the developed immunosensors have excellent specificity for the SARS-CoV-2 N-protein and little cross-reactivity for the MERS-CoV N-protein, influenza A virus N-protein, influenza B virus N-protein, and SARS-CoV-2 S-protein.

Saliva-Based Tests
To validate the detecting ability of the immunosensors for COVID-19 diagnostics in a more realistic scenario, the measurements were taken from commercial saliva to mimic the actual samples. Saliva is viscous and tends to congeal after collection, making it difficult to be accurately pipette for liquid-based manipulation. Therefore, PBS was used to mix with saliva to lower the viscosity. Previous studies showed that the SARS-CoV-2 N-protein could be directly detected in PBS-diluted saliva for COVID-19 rapid tests [11,39]. Following the PBS-based immunoassay, the ten times-diluted salivary solutions were spiked with SARS-CoV-2 N-protein. Figure 4 shows the ∆R et /R et0 results in the diluted saliva with and without N-protein. The linear regression equation of N-protein detection was ∆R et /R et0 (%) = 29.295 log[N-protein] (ng/mL) + 64.771 in the dynamic range of 0.01−100 ng/mL. The correlation coefficient was 0.991, implying good linearity. The calculated LOD was 6 pg/mL (S/N > 3), the same as the LOD obtained in PBS. The slope (29.30% mL/ng) of the calibration curve was slightly smaller than that (32.45%·mL/ng), attributed to the effect of the saliva viscosity on the immunoreaction efficiency [22]. Furthermore, the immunosensors were tested with the same immunoreacting procedures in SARS-CoV-2 Nprotein-free saliva and PBS, as shown in Figure 4a. The repetition test in blank solutions can realize the short-term stability of immunosensors and the effect of the interferents existing in saliva on the sensors. The ∆R et /R et0 values increased slightly with the repetition number measured in the N-protein-free saliva samples, and the slope of the linear regression curve was 1.644 (%·mL/ng), which was only 0.056 times smaller than that obtained in the Nprotein-containing saliva samples. The ∆R et /R et0 difference between the 0.01−100 ng/mL N-protein-containing saliva and the N-protein-free saliva was significant (p < 0.05) (t-test analysis). This result implies that the influence of a saliva substrate on the sensing result can be ignored. Compared to the N-protein-free saliva, the ∆R et /R et0 measured in blank PBS ranged from −0.1% to −1.7%, indicating that the sensor stability was good after repeatedly dripping blank PBS and mediator-containing PBS for the background test and the EIS measurements.
Furthermore, four healthy volunteers donated their saliva for real sample testing instead of patient saliva. First, the concentration-varied N-protein was spiked in the 10 times-diluted human salivary solutions, and then, the salivary solutions were filtered through 0.22-µm pore size syringe filters (PES membrane, Millex-GP, Merck Millipore Ltd., Dublin, Ireland). Some (50 µL) aliquots of the salivary filtrate were dripped on the immunosensors for 40 min, and then, PBS was used to rinse the immunosensors. Subsequently, the 2.5 mM Fe(CN) 6 3−/4 -containing PBS was dripped on the immunosensors for the EIS measurements. Figure 4b shows the ∆R et /R et0 signal obtained from three individual immunosensors, called S1, S2, and S3. The results showed that the three immunosensors had good linearity in 0.01−100 ng/mL. The LOD values obtained at the S1, S2, and S3 sensors were 6, 5, and 7 pg/mL. The results suggest that the constructed immunosensors are still feasible for real-saliva detection. Furthermore, the sensitivity of the S1-S3 sensor regression curves was smaller than that obtained in the diluted artificial saliva, attributed to the composition and viscosity of different volunteer saliva.
The methodology, preparing strategies, and sensing properties of the immunosensors are compared with other electrochemical affinity-based sensors in Table 2. Generally, the nanomaterial-modified immunosensors have promising potential to obtain a lower LOD [25,26,45]. Our BioAb/SA-BSA/MPA/SPCE-based immunosensor presented a low LOD, a wide linear range, and a simple label-free sensing strategy. Some studies and this work can reach pg/mL-scaled LOD [24][25][26]32,45]. Shan's clinical study found that, when the PCR Ct value was higher than 30, the corresponding N-protein concentration in saliva was lower than pg/mL [39]. Therefore, the future mission is to promote the LOD of immunosensors for more sensitive saliva-based COVID-19 antigen testing.  from 0.51% to 1.21% and from 0.58% to 1.84%, respectively; those are much small the ΔRet/Ret0 of SARS-CoV-2 N-protein. The signal-to-cross reactivity ratio of 1 ng/m 100 ng/mL SARS-CoV-2 N protein was larger than 64.2 and 79.5, respectively. The indicated that the developed immunosensors have excellent specificity for the SAR 2 N-protein and little cross-reactivity for the MERS-CoV N-protein, influenza A v protein, influenza B virus N-protein, and SARS-CoV-2 S-protein.

Saliva-Based Tests
To validate the detecting ability of the immunosensors for COVID-19 diagno a more realistic scenario, the measurements were taken from commercial saliva to the actual samples. Saliva is viscous and tends to congeal after collection, making cult to be accurately pipette for liquid-based manipulation. Therefore, PBS was u mix with saliva to lower the viscosity. Previous studies showed that the SARS-Co protein could be directly detected in PBS-diluted saliva for COVID-19 rapid tests Following the PBS-based immunoassay, the ten times-diluted salivary solution spiked with SARS-CoV-2 N-protein. Figure 4 shows the ΔRet/Ret0 results in the dilu liva with and without N-protein. The linear regression equation of N-protein de was ΔRet/Ret0 (%) = 29.295 log[N-protein] (ng/mL) + 64.771 in the dynamic range of 0. ng/mL. The correlation coefficient was 0.991, implying good linearity. The calculate was 6 pg/mL (S/N > 3), the same as the LOD obtained in PBS. The slope (29.30% m of the calibration curve was slightly smaller than that (32.45%•mL/ng), attributed effect of the saliva viscosity on the immunoreaction efficiency [22]. Furthermore, munosensors were tested with the same immunoreacting procedures in SARS-Co protein-free saliva and PBS, as shown in Figure 4a. The repetition test in blank so can realize the short-term stability of immunosensors and the effect of the inter existing in saliva on the sensors. The ΔRet/Ret0 values increased slightly with the rep   Furthermore, four healthy volunteers donated their saliva for real sample testing instead of patient saliva. First, the concentration-varied N-protein was spiked in the 10 times-diluted human salivary solutions, and then, the salivary solutions were filtered through 0.22-μm pore size syringe filters (PES membrane, Millex-GP, Merck Millipore Ltd., Dublin, Ireland). Some (50 μL) aliquots of the salivary filtrate were dripped on the immunosensors for 40 min, and then, PBS was used to rinse the immunosensors. Subse-

Conclusions
In this work, we constructed BioAb/SA-BSA/MPA/AuNS/SPCE-based immunosensors for the label-free detection of the SARS-CoV-2 N-protein in saliva samples. The MPA SAM for the SA-BSA immobilization allows EIS-based immunosensors to have a more sensitive response to the antigen-Ab immunoreaction than the MUA SAM. The oriented adsorption of BioAb on the SA-BSA layer can facilitate the sensitivity and reproducibility of the fabricated immunosensors. The immunosensors presented good linearity from 0.01 ng/mL to 100 ng/mL and a low LOD of 6 pg/mL in the diluted saliva. Moreover, the high signal-to-cross reactivity ratio of the immunosensors implies excellent specificity of the SARS-CoV-2 N-protein, indicating that the surface modification technique can effectively reduce the nonspecific adsorption of other viral antigens. The saliva-based immunosensor exhibits promising potential to develop a rapid and easy operational device for COVID-19 diagnostics.